The present disclosure relates to systems and methods for magnetic resonance imaging (“MRI”). More particularly, the disclosure relates to systems and methods for designing radio frequency (“RF”) pulses for in-vivo MRI applications, where the RF pulses are robust against errors due to physiological motion of organs during the respiratory cycle.
Cardiac and respiratory motion causes the human heart to be among the most challenging organs for MRI. However, research has made significant progress recently to addressing those and other challenges allowing cardiovascular magnetic resonance (CMR) to become an indispensable tool among the different techniques to diagnose cardiovascular diseases.
Cardiac motion is typically addressed by acquiring the data in synchrony with the cardiac cycle, enabled by an electrocardiogram, pulse oximeter, or acoustic triggering devices. Respiratory motion is addressed in a variety of ways. A large fraction of acquisitions is acquired under single or multiple breath-holds, either performed during full inhalation, which is often more pleasant for the volunteers, or during expiration, which typically results in lower respiration variability between different breath-holds. The other fraction of cardiovascular acquisitions is performed during free-breathing, because of long scan times, patient compliance, or because breathing is deliberately wanted to identify a disease.
Most free-breathing protocols include navigator scans into the sequence, which monitor the respiration level based on the position of the diaphragm. Accordingly, only scans within a predefined acceptance window (typically around 7 mm wide) are included in the reconstructions. Within the acceptance window, the slice position is often prospectively corrected according to the diaphragm position with a fixed factor of typically 0.6. Navigator based acquisitions are relatively inefficient since only 30-50% of the acquisitions commonly fall within the acceptance window.
In order to reduce scan times, several recent studies have demonstrated that scan efficiencies of up to 100% are feasible while data is corrected retrospectively for respiratory motion. Furthermore, recent advances in hardware, pulse sequence design, and reconstruction algorithms have pushed acquisition speed towards real-time cardiac imaging enabling image acquisition times of less than 50 ms and during this period respiratory motion can be neglected.
In addition to the above-mentioned advances in cardiac MRI, there is also an ongoing trend towards higher fields. Despite challenges of banding artifacts, higher specific absorption rate (SAR) and contrast non-uniformities, an increasing number of clinical scans are performed at 3T.
MR scanners operating at a higher main magnetic field strength (B0) provide higher signal-to-noise ratio (SNR) and better acceleration performances in parallel imaging techniques, allowing for higher spatial resolution images and/or shorter acquisition times. Higher magnetic fields also provide stronger tissue contrast in a variety of applications.
Today, most clinical MR scanners operate at a B0 field of 1.5 Tesla (T) or 3 T, with 1.5 T typically considered standard field and 3T considered “high field.” In recent years, a strong interest in systems operating at 7T, considered “ultra high field,” or UHF, resulted in several tens of human 7T systems being installed in academic research centers, with a growing body of clinical research studies published every year at 7T.
Despite gains in SNR and tissue contrast, increased main magnetic field strengths are also faced with several challenges, including magnetic susceptibility induced B0 inhomogeneities and inhomogeneities of the transmit B1 field, or radiofrequency coils. These two issues are complicated by physiological motion that can alter ΔB0 and B1+ maps.
With respect to magnetic susceptibility induced artifacts, when a human body is placed in the homogeneous B0 field of an MR scanner, spatial perturbations of B0 (ΔB0) will occur, which are mainly induced by different magnetic susceptibilities between different biological tissues of the human body. In the presence of large ΔB0, severe artifacts of multiple kinds typically occur in the resulting images. Because ΔB0 variations are proportional to B0, larger artifacts occur as the field increases. So-called B0 Shimming coils help countering ΔB0 variations by applying additional magnetic fields trying to cancel undesired ΔB0 within a given region of interest. However, B0 shimming can only achieve partial correction this problem, and the magnitude of residual artifacts increases as the main magnetic field B0 increases. The ΔB0 also affects the spatial excitation profile of the RF pulses utilized to excite spins to generate the MR signals that are then sampled by the receiver chain.
The second challenge associated with increasing field strength is the shortened wavelength of the transmit RF field, due to the fact that MR operates at the Larmor frequency of protons which is proportional to B0. This can lead to significant variations of the transmit magnetic field (B1+) magnitude which consequently results in spatial variations of image intensity and image contrast. This problem is especially significant in the torso where the ratio of RF wavelength over organ size is even smaller, such as in the liver or in the heart. At clinical field strength of 3T B1+ variations of more than 50% over the heart have been reported. At 7T field strength, B1+ variations are intrinsically stronger, and can even cause a complete loss of B1+ in local area. The resulting contrast and signal intensity variations can significantly affect scientific results and deteriorate the diagnostic quality of the MR images.
Addressing spatial inhomogeneity of RF excitation to restore homogeneous tissue contrast can be achieved using a transmission RF coil including multiple, independent transmitting coil elements, knowing that the final excitation B1+ field is the superposition of the complex B1+ fields of each coil element. The simplest method, referred to a “B1+ shimming” includes applying a constant complex factor on each coil element scaling the amplitude and phase of the input RF power of each coil element. The complex factors are optimized to obtain a homogeneous superposition of the individual B1 fields.
A more powerful and general approach, referred to as “parallel transmission,” or “pTX,” additionally includes temporal changes, which means that at each time point (typically every 4 to 10 microseconds) of an RF pulse (with typical duration of 0.5 to 4 ms), the complex RF input on each individual channel can be varied independently to the others. A technique often referred to as “spokes,” uses a series of sub-pulses (˜0.5 ms/subpulse), each being plaid with a specific B1+ shim solution. This temporal flexibility, together with the use of gradient encoding moments, greatly increases the degrees of freedom, with higher excitation fidelity, to the cost of higher complexity and more expansive hardware than required for B1+ shim.
The calculation of the complex B1+ shimming factors or the pTX RF pulses is based on calibration scans, acquired prior to the respective imaging scan. These typically include spatial mapping of the B1+ field of each transmit element and potentially a ΔB0 map. The calculated RF pulses or B1+ shim factors are then applied for a dedicated sequence during the imaging scan.
A significant challenge is the possibility that the respiratory status of the patient/subject changes between the calibration scans (B1+ and B0 maps) and the actual imaging scan for which the B1+ shim solution or pTX RF pulse are applied. In practice, a good number of scans in the torso are acquired during breath-holds, where the patient is asked to always come back to a same exhale (or inhale) position; however, patients will not always return to the same position. This patient motion can significantly impact resulting image quality (e.g., with degradation of excitation homogeneity). This deleterious impact primarily comes from significant differences between the B1+ maps of individual RF coil elements (more so than differences in B0 maps) at different phases of the respiration cycle.
Accordingly, systems and methods are needed to manage these competing imaging constraints and sources of potential deterioration of the quality of the resulting clinical images.